Except where specified below the term fibroin is used to refer generically to the main structural protein of cocoon silks whether they are derived from the domesticated Mulberry Silkworm (Bombyx mori) or a transgenic silkworm or from any Wild Silkworm including, but not limited to those producing Muga, Eri or Tussah silks. Furthermore, the term ‘silk’ is used to refer to the natural fine fibre that silkworms secrete, which comprises the two main proteins, sericin and fibroin, fibroin being the structural fibres in the silk, and sericin being the material surrounding the fibroin and sticking the fibres together in the cocoon. ‘Silk cocoon’ is used to refer to the casing of silk spun by the larvae of the silk worm for protection during the pupal stage.
Three types of cartilage are found in the body of mammals: white fibrocartilage; yellow elastic cartilage; and hyaline cartilage. Hereafter except where stated, the term cartilage is used in its generic sense including these three different types of cartilage.
White fibrocartilage is found in the menisci of the knee and the tempero-mandibular joint and in intervertebral discs. Yellow elastic cartilage is found in the pinna of the ear, the epiglottis and around the auditory canal. Hyaline cartilage is found mainly as articular cartilage in non-synovial joints where it provides smooth articulating surfaces and in synovial joints where it provides a hard and stiff connective tissue covering the articular surfaces of diarthroidal synovial joints.
Articular hyaline cartilage provides a long lasting, lubricated, low friction joint surface, distributes stresses over a broad area of underlying bone and may help to dissipate shocks during dynamic loading (Mow V C, Ratcliffe A. Structure and function of articular cartilage and meniscus. In: Mow V C, Hayes W C, eds. Basic Orthopaedic Biomechanics. New York: Raven Press, 1991; 143-198). The compressive stiffness of cartilage is extremely important in its function. The stiffness of viscoelastic materials such as cartilage depends greatly on the loading history of the material and the method for measuring it and consequently several moduli are used to describe articular cartilage. For example, Spiller, K. L., Laurencin, S. J., Charlton D, Maher, S. A., Lowman, A. M. (2008) in their paper “Superporous hydrogels for cartilage repair: Evaluation of the morphological and mechanical properties” Acta Biomaterialia 4, 17-25, state that the unconfined compressive elastic modulus of adult articular cartilage is about 1 MegaPascal (MPa), while the aggregate compressive modulus is about 0.33 MPa. Treppo, S. et al. in Comparison of biomechanical and biochemical properties of cartilage from human knee and ankle pairs, Journal of Orthopaedic Research 18, 739-748 (2000), state that the equilibrium modulus of healthy adult human articular cartilage lies between 0.2 and 1.5 MPa with a mean about 0.6 MPa depending on, which joint the cartilage is taken from, the location on the joint and depth. Park, S., Hung, C. T. & Ateshian, G. A. in Mechanical response of bovine articular cartilage under dynamic unconfined compression loading at physiological stress levels, Osteoarthritis and Cartilage 12, 65-73 (2004), state that the unconfined dynamic modulus for bovine tibial cartilage lies between 15-65 MPa depending on applied stress and loading frequency.
The menisci of the knee joint are crescent shaped discs, largely constructed from white fibrocartilage. They are interposed between the femoral condoyle and tibial plateau and have the function of compressive load spreading, shock absorption, stabilization and secretion of synovial fluid for articular lubrication. The structure, function and pathology of the menisci have been reviewed by S. M. Bahgia and M. Weinick, Y. Xing, and K. Gupta (2005) Meniscal Injury, E-medicine World Library, 27 Jul. 2005, www.emedicine.com/pmr/topic75.htm. The outer rim is vascular while the central part is avascular fibrocartilage. The menisci contain 70% type I collagen (non-articular cartilage fibrillar collagen). The collagen fibres of the meniscus show a predominantly circumferential orientation together with some radial tie fibres. Collagen orientation is extremely important for the mechanical function and fixation of this structure. Compression of the meniscus leads to tensile hoop loading of the circumferential fibres and radial loading of the radial fibres, resisting spreading and flexing of the menisci. Thus the ability of the meniscus to spread load and dissipate energy is dependent on the integrity of the collagen fibre lay. For this reason damage to these fibres increases the risk of secondary osteoarthrotic damage to the condylar cartilages as the normal load distribution and shock-absorbing functions are impaired.
Meniscal injuries are fairly common in adults and are frequently sports-related. They are less common in children over 10 years old and rare in children under 10 with morphologically normal menisci (Iobst, C. A. and Stanitski, C. L., 2000, Acute knee injuries. Clin Sports Med. 2000 October; 19(4):621-35).
Total knee replacement involves the insertion of a highly complex metal and polymer implant and cannot be considered as treatment for uncomplicated meniscal injury. The Dacron and Teflon meniscal component may initiate severe synovial reactions (Cook, J. L., Tomlinson, J. L., Kreeger, J. M., and Cook, C. R. 1999. The American Journal of Sports Medicine 27:658-665 Induction of meniscal regeneration in dogs using a novel biomaterial) while loosening and mechanical failure are a problem (de Groot, J. H. 1995 Doctoral dissertation. University of Gronigen, Summary p153).
Surgical treatment of damaged menisci is often necessary, for which there are different surgical treatment options.
Small meniscal tears can be repaired directly using sutures, fasteners or arrows. However small tears account for less than <3% of all presented mensical injuries.
Although total or partial removal of the meniscus (meniscectomy) to remove damaged meniscal tissue was popular some forty years ago, it is well understood that this procedure leads to articular cartilage degeneration (King, D. Clin. Orthop. 1990, 252, 4-7; Fairbank, T. J. Journal of Bone and Joint Surgery 1948, 30, 664-670) in turn leading to osteoarthrosis. The extent of the secondary osteoarthrosis caused by menisectomy appears to depend on how much meniscal tissue has been removed. Therefore partial meniscectomy usually involving the removal of about 25-40% of the meniscal tissue is the current most frequently used procedure. However, even with partial mensicectomy, a reduction in both shock absorption and the stability of the knee results in secondary osteoarthrosis in the medium to long-term. Better alternatives to partial meniscectomy are therefore being sort. Allograft transplantation is only partially successful as an alternative to total or partial menisectomy so currently only about 0.1% of meniscal procedures employ this approach. There is no proof that replacement of the meniscus with an allograft can re-establish some of the important meniscal functions, and thereby prevent or reduce the development of osteoarthrosis secondary to meniscectomy (Messner, K. and Gao, J. 1998. The menisci of the knee joint. Anatomical and functional characteristics, and a rationale for clinical treatment. Journal of Anatomy, 193:161-178). The major problems are the lack of remodeling of the graft resulting in inferior structural, biochemical and mechanical properties and insufficient fixation to bone (Messner and Gao 1998, Loc. cit). Further disadvantages include the shortage of suitable donors, difficulties with preservation techniques, the possible transfer of diseases, difficulty in shaping the implant to fit the donor and possible immunological reactions to the implant (Stone, K. R. Clinical Sports Medicine. 1996, 15: 557-571).
In addition to allograft procedures, a number of implantable materials have been suggested as replacements for surgically removed damaged meniscal tissue. These include: collagen treated with pepsin to render it substantially non-immunogenic and subsequently cross-linked with glutaraldehyde; a material made from the submucosa of the small intestine; cross-linked hyaluronic acid, Teflon fibre; carbon fibre; reinforced polyester; and polyurethane-coated Dacron. The mechanical properties of these implant materials are a poor match for those of meniscal fibrocartilage which has an unconfined compressive elastic modulus of about 0.4 to 0.8 MPa. These materials have poor resistance to wear and are not self healing. Some of the above are non-resorbable, and are not replaced by functional tissue in situ. It is therefore not surprising that partial or total meniscal replacements made from collagen, Teflon fibre, carbon fibre, reinforced polyester, or polyurethane-coated Dacron showed high mechanical failure rates (de Groot 1995 loc. cit.). Failure also results from poor fixation and severe inflammatory response (de Groot 1995 loc. cit.).
Elastomers based on amphiphilic urethane block copolymers have been suggested for meniscal repair and tested in an animal model. (Heijkants, R. G. J. C. 2004 Polyurethane scaffolds as meniscus reconstruction materials, Ph.D. Thesis, University of Groningen, The Netherlands, MSC Ph.D.-thesis series 2004-09; ISSN: 1570-1530; ISBN: 90 367 2169 5, chapter 10 pp 167-184). These materials are likely to produce less toxic degradation products than Dacron or Teflon. However, the mechanical properties of the polyurethanes tested did not match native meniscus very well (Heijkants 2004 loc. cit.) and this may help to explain why only poorly orientated collagen was found in the regenerating fibrocartilage in the implanted devices in place of the well-orientated collagen in a normal meniscus. A further potential problem was that the polyurethane materials produced a Stage I inflammatory response (giant cells and some macrophages) (Heijkants 2004 loc. cit.). A follow up study tested a polycaprolactone-polyurethane co-polymer porous meniscal repair device over a two year period. After the testing period the device demonstrated no resorption capability, was not replaced by functional meniscal tissue and demonstrated no prevention of cartilage damage (Welsing R. T. C, van Tienen, T. C., Ramrattan, N., Heijkants, R., Schouten, A. J., Veth, R. P. H. and Buma, P. 2008; Effect on tissue differentiation and articular cartilage degeneration of a polymer meniscal implant: a 2 year follow up study in dogs. Am. Jour. Sports Med. 36 1978-1989).
Recently, tissue engineering strategies for meniscal repair have been suggested including the use of biocompatible scaffolds as a substrate for regeneration, and cellular supplementation to promote remodeling and healing. Little is known, however, about the contributions of these novel repair strategies to the restoration of normal meniscal function (Setton, L. A., Guilak, F, Hsu, E. W. Vail, T. P. 1999 Biomechanical Factors in Tissue Engineered Meniscal Repair. Clinical Orthopaedics & Related Research. (367S) supplement: S254-S272, October 1999).
Intervertebral discs lie between the cartilage end caps covering the ends of the vertebral centra. They consist of an outer annulus fibrosus, which surrounds the inner nucleus pulposus. The annulus fibrosus consists of several layers of fibrocartilage. The nucleus pulposus contains loose collagen fibrils and chondrocytes suspended in a mucoprotein gel. Intervetebral discs provide a deformable space between the vertebral bodies which facilitates flexibility of the vertebral column while at the same time acting as a shock absorber (M. D. Humzah And R. W. Soames 1988 “Human Intervertebral Disc: Structure And Function”, The Anatomical Record 220:337-356). Prosthetic discs are used to replace damaged discs in patients with herniated lumbar intervertebral discs, degenerative disc disease in the lumbar region, or post-laminectomy syndrome. They are also used to treat patients with lower back pain refractory to conservative treatment for more than six months and patients currently considered suitable for spinal fusion surgery (NICE guidelines, www.nice.org.uk/guidance/index.jsp?action=byID&r=true &o=11081).
There are significant problems associated with the use of metal-containing and non-metallic prostheses for total disc replacement.
Resilience is an extremely important property for natural meniscal and articular cartilage and for materials used to repair them. Resilience can be defined as the extent to which the material returns to its original thickness after being compressed. More precisely it can be defined as the property of a material to store energy reversibly when it is deformed elastically. In the context of articular and meniscal cartilage it is important as it is a measure of the ability of the material to recover from the deformation caused by the compressive loading produced by standing, walking, running and other movements. The high resilience of meniscal cartilage is also important as it enables it to function as an efficient shock absorber during the repeated loading cycles of walking and running. Resilience can be measured in a number of different ways. Most accurately resilience is the maximum energy per volume that can be elastically stored and is therefore measured by determining the area under the elastic part of the stress-strain curve. The resilience of human articular cartilage measured in this way gave a value of 2.9 Jm−3 (Park, S. S., Chi, D. H., Lee, A. S., Taylor, S. R. & Iezzoni 2002, J. C. “Biomechanical properties of tissue-engineered cartilage from human and rabbit chondrocytes” Otolaryngology and head and neck surgery 126, 52-57). However it is simpler to use a measure of the extent to which the deformation is recoverable after one or more loading cycles.
Destruction of the articular cartilage on the articular surfaces resulting in changes to the bone adjacent to the articular cartilage occurs in the condition osteoarthrosis. This commonly affects hip, knee, hands, feet and spinal joints. It causes chronic pain, loss of mobility and often stiffness. Primary osteoarthrosis is mainly an effect of the aging process while secondary osteoarthrosis is caused by changes in the stress distribution in joints resulting from injuries, obesity, ligament degeneration, hardness of the subchondral bone or genetic changes to the joint morphology. Disease states such as diabetes mellitus and gout, and other factors including hormonal changes are also causal in secondary osteoarthrosis. The disease process in primary and secondary osteoarthrosis is the same. Severe osteoarthrosis is commonly treated by insertion of an artificial hip or knee joint prosthesis made from a range of materials including alloy steel, ceramic and synthetic polymers. However, these procedures are expensive and not without risk. In addition, about 4% of total hip replacements fail within 10 years as a result of aseptic loosening, deep infection, prosthesis fracture and other causes. Total knee prostheses in general use have an average rate of failure of about 1% per year while those types of prosthesis less frequently used show failure rates up to 3 times faster. More than one complicated and expensive revision may therefore be required in the lifetime of a young person receiving a total joint replacement. The toxicity of polymeric and metallic wear products is an additional problem. Thus much work has been carried out to find viable alternative procedures to joint replacement for the repair of articular cartilage and hence the prevention and treatment of osteoarthrosis.
Suggested methods for articular cartilage repair can be put into two categories; cell-based and tissue-based methods.
Cell-based methods can be further divided according to whether the cells are implanted with or without incorporation into a matrix and whether the matrix is biodegradable or non-biodegradable. Cell based methods include marrow stimulating techniques, autologous chondrocyte implantation, matrix assisted autologous chondrocyte implantation, and procedures using expanded mesenchymal stem cell cultures. For small focal cartilage lesions cell-based cartilage repair strategies can give better clinical outcomes compared with no treatment. However cell-based strategies cannot be used to treat severe osteoarthrosis and this approach has other problems and limitations (Richter, W. 2007. Cell-based cartilage repair: illusion or solution for osteoarthritis. Current Opinion in Rheumatology 19 (5) 451-456).
Tissue-based methods for articular cartilage repair include autologous perichondrial, periosteal or osteochondral grafts and the implantation of allogenous osteochondral and chondral grafts. There are significant problems and limitations associated with tissue-based methods for articular cartilage repair. For example, autologous osteochondral grafts involve the potentially damaging removal of healthy osteochondral tissue from the joint while problems of allograft procedures include the scarcity of fresh donor material, damage to the graft from immune attack, mechanical deterioration and death of chondrocytes during graft handling and frozen storage and risk of disease transmission. (Hunziker, E. B. 2001 Articular cartilage repair: basic science and clinical progress. A review of the current status and prospects can be found in Osteoarthritis and Cartilage (2001) 10, 432-463).
Scaffolds for articular cartilage replacement whether implanted on their own or containing cells, must have appropriate porosity to allow for cell migration and nutrition and mechanical properties similar to those of healthy cartilage to enable load bearing and provide a tough, stiff, low friction articular surface as stated by S. Frenkel and P. D. Cesare, Scaffolds for articular cartilage repair, Annals of Biomedical Engineering 32 (1) (2004), pp. 26-34. and by S. J. Hollister, Porous scaffold design for tissue engineering, Nature Materials (7) (2006), p. 590. In addition, the scaffold should be capable of rapid remodeling (Hunziker, E. B. (1999) Articular cartilage repair: are the intrinsic biological constraints undermining this process insuperable? Osteoarthritis and Cartilage 7, 15-28). Existing methods for repairing articular cartilage suffer from the generic problems of poor mechanical properties, poor tissue integration and chondrocyte loss from the lesion borders while the first of these is the most serious (Hunziker 1999 op. cit.). Scaffolds prepared from synthetic biodegradable polymers such as poly(D,L-lactic-coglycolic acid) (PLGA) generally have good mechanical properties, but are reabsorbed too quickly to give sufficient time for the formation of new tissue. In addition, they generally bind cells poorly as a result of their hydrophobic surfaces and lack of cell adhesion signals. In contrast biological macromolecules generally show better cell binding but poor mechanical properties: Janaár, J. Slovíkovä, A. Amler, E., Krupa, P., et al. Mechanical Response of Porous Scaffolds for Cartilage Engineering. Physiological. Research. 56 (Suppl. 1): S17-S25, 2007. Existing degradable scaffolds are generally too weak to support the forces found in load bearing cartilage. Spiller, K. L., Laurencin, S. J., Charlton D, Maher, S. A., Lowman, A. M. (2008) Superporous hydrogels for cartilage repair: Evaluation of the morphological and mechanical properties Acta Biomaterialia 4, 17-25. These authors teach the use of non-degradable hydrogel scaffolds prepared from a mixture of the non-degradable polymers poly(vinyl alcohol) and poly(vinyl pyrolidone) incorporating degradable microparticles of poly(lactic-co-glycolic acid). The resulting scaffolds were considerably less stiff (unconfined compressive elastic modulus up to 0.15 MPa; aggregate compressive modulus up to 0.14 MPa) compared with the comparable values of 1 MPa and 0.33 MPa respectively for adult articular cartilage. Thus these scaffolds are neither biodegradable nor comparable to articular cartilage in compressive properties.
U.S. Pat. No. 6,306,169 discloses an implant with a porous macrostructure infiltrated with hydrated polymeric gel. The structure is made from a bioresorbable polymer (poly-L-lactic acid, polycaprolactone, polyhydroxybutarate, or polyanhydrides) and the gel comprises alginate, agarose, carrageenans, glycosaminoglycans, proteoglycans, polyethyelene oxide or collagen monomers.
U.S. Pat. No. 6,514,515, U.S. Pat. No. 6,867,247 and U.S. Pat. No. 7,268,205 disclose a bioresorbable and biocompatible polymer, polyhydroxyalkanoate, for a range of implantable applications including the repair of meniscal and articular cartilage.
De Groot, “Meniscal tissue regeneration in porous 50/50 copoly(L-lactide(epsilon-caprolactone) implants,” Biomaterials 18(8):613-22 (1997) discloses the use of porous copoly(L-lactide(epsilon-caprolactone) for meniscal tissue regeneration.
U.S. Pat. No. 6,747,121 discloses the use of a porous resorbable implantable material comprised of a terpopolymer containing L-lactide, a glycolide and one other type of repeat unit selected from the group consisting of D-lactide, D,L-lactide and epsilon-caprolactone.
U.S. Pat. No. 6,103,255 teaches the use of biocompatible and biodegradable polymers for use as components of tissue scaffolds. Such polymers include polycarbonates, polyarylates, block copolymers of polycarbonates with poly(alkylene oxides), block copolymers of polyarylates with poly(alkylene oxides), a-hydroxycarboxylic acids, poly(capro-lactones), poly(hydroxybutyrates), polyanhydrides, poly(ortho esters), polyesters and bisphenol-A based poly(phosphoesters).
U.S. Pat. No. 6,679,914 discloses a meniscal prosthesis comprising a plurality of superimposed sheets of animal pericardium cross-linked by an aldehyde. Although the device is likely to be resorbable and may be biocompatible despite the use of aldehyde cross-linking, the patent does not disclose the mechanical properties of the device, which are likely to be considerably inferior to those of the normal meniscus.
CA 2,374,169 discloses a biocompatible, resorbable implantable material for total replacement or reinforcement of connective tissue. The material comprises a flexible elongate tape and a plurality of aligned fibres, the tape comprising two essentially parallel layers of mesh and a hydrogel. The material can be of poly(lactic acid), poly(glycolic acid), polydioxanone, polycaprolactone, polyhydroxybutyrate, poly(trimethylene carbonate) or mixtures of these materials.
Several additional problems and limitations have been noted in synthetic polymer scaffolds in addition to their generally poor match in compressive properties compared with those of cartilage. Polymers containing lactic and/or glycolic acids have been shown to give rise to toxic solutions probably as a result of acidic degradation as described by Tayler M S, Daniels A U, Andriano K P, Heller J (1994) Six bioabsorbable polymers: In vitro acute toxicity of accumulated degradation products. Journal of Applied Biomaterials 5: 151-157. This is of particular concern in connection with cartilage repair in which relatively large quantities of synthetic polymer may be required and where poor vascularity slows the removal of toxic waste products. In addition to lactic and glycolic acid many other biodegradable synthetic polymers contain acidic units including butyric, valeric and caproic acids and it is possible that acidic breakdown products from these may also be toxic. A further concern with poly(lactic) and poly(glycolic) acid is that small particles arise during degradation and these can trigger an inflammatory response as reported by Gibbons D F (1992) “Tissue response to resorbable synthetic polymers”. In Degradation Phenomena on Polymeric Biomaterials, Plank H, Dauner M, Renardy M, eds. Springer Verlag, New York. pp 97-104., extensive foreign-body response and osteolytic reactions have been reported in an orthopaedic use of polyglycolic acid as reported by Böstman O, Partio E, Hirvensalo E, Rokannen P (1992), Foreign-body reactions to polyglycolide screws, Acta Orthop Scand 63: 173-176. Similar responses are seen with poly(lactide) as reported by Bergsma E. J., Brujn W, Rozema F. R., Bos R. M., Boering G., Late tissue response to poly(L-lactide) bone plates and screws, Biomaterials 1995; 16(1):25}31. Although biodegradable polyurethanes appear to be satisfactory in in vitro and in vivo trials, urethane monomers are carcinogenic and the long effect of their degradation products and how those products are removed from the body is not clearly understood Gunatillake, P. A. and Adhikari, R. 2003, Biodegradable synthetic polymers for tissue engineering, European Cells and Materials, 5.1-16.
WO 2005/094911 discloses a composite material comprising one or more silk elements in an acrylic or cross-linked protein matrix. The material can be used in a wide range of implantable devices and can be made from certain Wild silks naturally decorated with the integrin-binding tripeptide RGD. This tripeptide in Wild silks may facilitate the binding of mesenchymal and other cells. The material was prepared according to the standard protocol described for example by Chen, X., Knight, D. P., Shao, Z. Z., and Vollrath, F. (2001) “Regenerated Bombyx silk solutions studied with rheometry and FTIR” Polymer, 42, 9969-9974. The document reports that the standard protocol results in considerable degradation of the fibroin, which would yield scaffolds with reduced strength, stiffness and resilience.
The standard protocol for preparing regenerated fibroin solutions involves degumming in hot alkaline solutions and dissolution in hot 9M to 9.5M lithium bromide solution. Recently, it has been suggested that cartilage-like materials prepared in vitro by culturing sponges made of regenerated silk fibroin inoculated with chondrocytes or mesenchymal stem cells may have potential for cartilage repair as reviewed by Hofmann S, Knecht S, Langer R, Kaplan D L, Vunjak-Novakovic G, Merkle H P, Meinel L: “Cartilage-like tissue engineering using silk scaffolds and mesenchymal stem cells”. Tissue Engineering 2006, 12(10):2729-2738 and by Vepari C, Kaplan DL: “Silk as a biomaterial”. Progress in Polymer Science (Oxford) 2007, 32(8-9):991-1007. Little work however appears to have been done to characterize the mechanical properties of fibroin scaffolds which appear to be considerably less stiff, less strong and with higher friction surfaces than adult articular cartilage. However, three papers by Morita's group describe an initial attempt to define the effect of time in culture on the compressive properties of a potential cartilage replacement material grown in vitro by seeding a porous fibroin sponge with chondrocytes. This group showed that the dynamic compressive modulus in this material decreased with time while creep deformation increased with longer cultivation as disclosed by Morita Y, Ikeuchi K, Tomita N, Aoki H, Suguro T, Wakitani S, Tamada Y: “Evaluation of dynamic visco-elastic properties during cartilage regenerating process in vitro”. Bio-medical materials and engineering 2003, 13(4):345-353 and the same authors in “Visco-elastic properties of cartilage tissue regenerated with fibroin sponge”, Bio-Medical Materials and Engineering 2002, 12(3):291-298. In addition, the coefficient of friction of the surface of a potential cartilage replacement material grown in vitro by seeding a porous fibroin sponge with chondrocytes was initially as low as that of natural cartilage but increased with increasing duration of a sliding test as a result of exudation of interstitial water from the surface layer as disclosed by Morita Y, Tomita N, Aoki H, Sonobe M, Suguro T, Wakitani S, Tamada Y, Ikeuchi K in their paper entitled, “Frictional properties of regenerated cartilage in vitro”. Journal of Biomechanics 2006, 39(1):103-109.
WO 2007/020449 discloses an implantable cartilaginous meniscal repair device partly or wholly comprised of porous fibroin. The regenerated fibroin used was prepared using the standard protocol resulting in a scaffold with reduced strength, stiffness and resilience.
WO 2004/US00255, US 20040107 and US 2007/0187862 disclose methods for producing porous silk fibroin scaffold material. A regenerated silk fibroin is first prepared converted to a porous fibroin scaffold using either salt leaching or gas foaming followed in both cases by treatment with methanol or propanol to stiffen and strengthen the material. The material is intended for use as a scaffold for growing a cartilage-like material in vitro. The protocol used for the preparation of the fibroin solution described in US 2007/0187862 involves degumming by boiling cocoons for 20 minutes in an aqueous solution of 0.02 M sodium carbonate solution followed by dissolution of the fibroin in 9.3 M lithium bromide at 60° C. Thus the protocol they used is closely similar to the standard protocol described in the literature and to that used by Holland, C., Terry, A. E., Porter, D. & Vollrath, F. in their paper “Natural and unnatural silks”, Polymer 48, 3388-3392 (2007). The latter authors prepared a regenerated fibroin solution prepared according to the standard protocol and compared the rheology of this and native Bombyx mori silk fibroin solution taken directly from the silk gland of the silkworm and at the same protein concentration. They concluded that the vast reduction of viscosities and storage modulus values they observed in the regenerated silk fibroin could be explained by degradation of both molecular weight and folding of the fibroin as a consequence of the protocol used. Thus there is strong evidence that the conditions for preparing the fibroin solution disclosed in US 2007/0187862 produce a marked degradation of the fibroin. This is likely to have a markedly negative impact on the compressive strength, moduli and resilience of the porous material produced from the silk fibroin solution.
WO 2007/020449 teaches an implantable cartilage repair device comprised of a three dimensional biomimetic fibrelay and a bioresorbable porous hydrogel. The hydrogel can be at least partially comprised of regenerated fibroin. The fibroin for the preparation of the porous hydrogel is prepared using the standard protocol comprising the steps of dissolving degummed silk in hot 9.3M lithium bromide solution; dialyzing resulting solution exhaustively against deionised water for two days and concentrating it in a vacuum dessicator.
US 2005/0281859 describes a method of forming an object from a feedstock, such as fibroin, capable of undergoing a sol-gel transition by adjusting the conditions to cause the feedstock to flow and then adjusting the conditions to gel the feedstock.
It has recently been shown that porous fibroin hydrogels prepared from the standard protocol are weak and have reduced resilience. Thus, there is still scope for improvement in the implantable materials and implants used for the replacement, partial replacement, or augmentation, or repair of damaged cartilage.
It is therefore, an object of the present invention to provide an improved regenerated fibroin solution and method of preparing an improved regenerated fibroin solution.
Another object of the invention is to provide an implantable fibroin material and a method of preparing the fibroin material, having improved mechanical properties.
It is a further object of the invention to provide an implant for the total or partial replacement, augmentation or repair of cartilage.